Magnetic resonance imaging (MRI) is commonly used to image the internal tissues of a subject. MRI is typically performed by placing the subject or object to be imaged at or near the isocenter of a strong, uniform magnetic field known as the main magnetic field. The main magnetic field causes the atomic nuclei that possess a magnetic moment in the matter comprising the subject or object to become aligned in the magnetic field. The nuclei (spins) begin a precession around the magnetic field direction at a rate which is proportional to the magnetic field strength. For hydrogen nuclei (which are the common nuclei employed in MRI), the precession frequency is approximately 64 MHz in a magnetic field of 1.5 Tesla.
In addition to the main magnetic field, a magnetic field gradient is also applied to form a magnetic resonance image. The magnetic field gradient in an MRI scanner is linear only within a limited region near the magnet isocenter. Outside of this region, virtually all gradient systems display nonlinear spatial characteristics, particularly at or near the edge of the magnet. This non-ideal condition is exacerbated by a rapid change of the main magnetic field (denoted as B0-field) toward the end of the magnet bore. As a result, the overall magnetic field produced by the combination of the gradient field and the B0-field has a complicated spatial dependence. In regions away from the isocenter, the overall magnetic field experienced by spins can be equal to the net magnetic field at or near the magnet isocenter. (Steckner et al., 1995, ISMRM Abstracts, pg. 756, Nice, France). These regions, sometimes referred to as the “gradient null”, are typically outside the imaging volume of interest. (King et al., U.S. Pat. No. 7,250,762). When a radiofrequency (RF) coil (or a coil element in a phased array) receives signals from that region, the signal will carry the same or similar frequency as the signal near the isocenter, leading to an aliasing artifact in the image. In a fast spin echo (FSE) pulse sequence, the aliasing artifact manifests itself as a series of spots, a band, or a “featherlike” artifact at or near the center of the field of view (FOV) along the phase-encoding direction. The artifact is often observed on sagittal or coronal planes in spine and knee scans, and can interfere with image interpretation. (Kim et al., 1999, ISMRM Abstracts pg. 1033, Philadelphia, Pa.). Various terms have been used to identify this artifact including, for example, cusp artifact, annefact, fold-over artifact, feather artifact, and peripheral signal artifact, along with other names. See, the Steckner, King, and Kim references, ibid. Although this artifact does not appear in exactly the same form (i.e., it is sometimes “C”-shaped), the mechanism of the artifact formation remains substantially the same.
One technique designed to reduce this artifact relies on adaptive phased-array coils (Frederick and Johnson, U.S. Pat. No. 6,134,465). Individual elements of a phased-array coil can be chosen automatically by an algorithm that determines the proper coil elements based on user-specified FOV, while also rejecting the signals from coil elements at or near the artifact-prone regions. However, this approach has been shown to be effective only under specific conditions (e.g., imaging with limited FOV). Further, in order to implement adaptive phased array, substantial modifications to the RF receiving electronics are required in addition to a signal selection algorithm.
To avoid excessive costs associated with hardware modifications, signal-processing techniques based on parallel imaging (e.g., sensitivity encoding) have been used to reduce the FSE cusp artifact. See, U.S. Pat. No. 7,250,762; Larkman et al., 2000, J Magn Reson Imaging 12:795-797; Pruesssmann et al., 1999, Magn Reson Med, 42:952-962. These signal-processing techniques are based on estimating an amplitude of the RF field sensitivity matrix by utilizing two separate coils, one placed at the magnet isocenter and the other (typically smaller in size) at or near the artifact-producing region. The non-aliased signal within the FOV can be recovered using a parallel-imaging reconstruction algorithm (Pruessmann, Id.). However, this approach requires knowledge of the approximate location of the artifact-producing region and also requires a calibration procedure to estimate the sensitivity matrix for each RF coil. These limitations can impose problems in practical implementation of the method.
Another approach to reducing such an artifact utilizes a metal foil (also known as “metal skirt” or “RF blanket”) over the artifact-producing region in order to dephase the magnetization leading to the artifacts. To be effective, this technique needs the RF blanket to be positioned exactly at the location of the artifact source (i.e., the precise location of the artifact source must be known). This method also raises safety concerns due to the possibility of increased local heating. (Schaefer, 1998, Magn Reson Imaging Clin N Amer. 6:775-789). These safety concerns can become prohibitive in a SAR-intensive sequence, such as FSE, particularly at high magnetic fields (e.g., 3.0 T).
Accordingly, there is a need in the field of MRI for straightforward, cost effective, and safe techniques for reducing the cusp artifact.